The invention relates generally to x-ray detectors and, more particularly, to an apparatus and method of reducing scattered x-ray detection in an x-ray detector.
X-ray imaging is a non-invasive technique to capture images of medical patients for clinical diagnosis as well as inspect the contents of sealed containers, such as luggage, packages, and other parcels. To capture these images, an x-ray source irradiates a scan subject (or object) with a fan beam of x-rays. The x-rays are then attenuated as they pass through the scan subject. The degree of attenuation varies across the scan subject as a result of variances in the internal composition of the subject. The attenuated energy impinges upon an x-ray detector designed to convert the attenuating energy to a form usable in image reconstruction. A control system reads out electrical charge stored in the x-ray detector and generates a corresponding image. For a conventional, screen film detector, the image is developed on a film and displayed using a backlight.
Presently, flat panel, digital x-ray detectors are being used to acquire data for image reconstruction. Flat panel detectors are generally constructed as having a scintillator, which is used to convert x-rays to visible light that can be detected by a photosensitive layer. The photosensitive layer includes an array of photosensitive or detection elements, where each element stores electrical charge in proportion to the light that is individually detected. Generally, each detection element has a light sensitive region and a region of electronics to control the storage and output of electrical charge. The light sensitive region typically includes a photoconductor, and electrons are released in the photoconductor when exposed to visible light. During this exposure, charge is collected in each detector element and is stored in a capacitor situated in the electronics region. After exposure, the charge in each detector element is read out using logic controlled electronics.
Each detector element may be controlled using a transistor-based switch. In this regard, the source of the transistor is connected to the capacitor, the drain of the transistor is connected to a readout line, and the gate of the transistor is connected to a scan control interface disposed on the electronics in the detector. When negative voltage is applied to the gate, the switch is driven to an OFF state, i.e., no conduction between the source and drain. On the other hand, when a positive voltage is applied to the gate, the switch is turned ON resulting in connection of the source to the drain. Often, each detector element of the detector array is constructed with a respective transistor and is controlled in a manner consistent with that described below.
For example, during exposure to x-rays, negative voltage is applied to all gate lines resulting in all the transistor switches being driven to or placed in an OFF state. As a result, any charge accumulated during exposure is stored in each detector element capacitor. During read out, positive voltage is sequentially applied to each gate line, one gate at a time. In this regard, generally only one detector element is read out at a time. A multiplexer may also be used to support readout of the detector elements in a raster fashion. An advantage of sequentially reading out each detector element individually is that the charge from one detector element does not pass through any other detector elements. The output of each detector element is then input to a digitizer that digitizes the acquired signals for subsequent image reconstruction on a per pixel basis. Each pixel of the reconstructed image corresponds to a single detector element of the detector array.
As described above, digital x-ray detectors utilize a layer of scintillating material, such as Cesium iodide (CsI), to convert incident radiation to visible light that is detected by light sensitive regions of individual detector elements of a detector array. Generally, transistor controlled detector elements are supported on a thin substrate of glass. The substrate, which supports the detector elements as well as the scintillator layer, is supported by a panel support. The panel support is not only designed to support the detector components, but also isolates the electronics that control the detector from the image detecting components. The electronics are supported by the panel support and enclosed by the back cover.
Many x-ray systems employ an anti-scatter grid. A primary function of the anti-scatter grid is to preferentially pass primary x-rays and reject scattered x-rays (e.g., Compton scattered x-rays). Accordingly, unwanted x-ray scatter is generally reduced when an anti-scatter grid is employed.
Scattered x-rays are objectionable because they often cause noise (e.g., image artifacts) to be present in a resulting x-ray image. Absent an anti-scatter grid, an x-ray image is often degraded by x-rays that are Compton scattered through the patient or object. Such scattered x-rays generally blur resulting images. As such, clinicians such as radiologists may have a difficult time interpreting such degraded images.
Typically, an anti-scatter grid is freestanding and is moveable relative to an x-ray detector. Such an anti-scatter grid is placed outside an x-ray detector. Often, such grids are packaged in a “cassette” located just beneath a patient support plate or Bucky cover. The cassettes generally require grid covers that provide mechanical support. These grid covers along with the Bucky cover, however, are known to absorb x-rays. Since these covers absorb x-rays, it is often necessary to increase the x-ray dosage to obtain desired image properties.
An anti-scatter grid typically includes a plurality of thin strips of a highly x-ray absorbing material, such as lead, that are placed parallel to one another along an edge. Rather than overlapping, however, the strips, generally referred to as septa, are typically arranged such that a uniform gap appears between edges of adjacent septa. Further, the septa are angled relative to one another such that the anti-scatter grid is substantially focused toward the x-ray source. As such, non-scattered x-rays from the x-ray source are more likely to pass through the gaps of the anti-scatter grid, and scattered x-rays are more likely to be absorbed by the anti-scatter grid. Accordingly, such anti-scatter grids are designed, in theory, to allow primary x-rays to pass therethrough while absorbing scattered x-rays. Unfortunately, in practice, such anti-scatter grids also absorb some primary x-rays. Typically, there is a tradeoff between good scatter rejection (absorption) and high primary x-ray transmission. This tradeoff is often captured in the quantum improvement factor (QIF). Typically, QIF Values above 1 indicate better imaging while those below 1 indicate that the anti-scatter grid may be degrading image quality. A typical mammographic anti-scatter grid might have a QIF value of about 1.05 to 1.1.
For x-ray detectors having a very fine pixel structure, such as detectors employed in mammography, the grid septa pattern may interfere with the pixel pattern on the detector. Such interference often manifests itself as interference lines, commonly called Moire patterning, in the resulting x-ray image. In such cases, the anti-scatter grid is moved relative to the x-ray detector during x-ray exposure to eliminate the patterning. Unfortunately, increased costs and complexities are generally associated with anti-scatter grids having moving capabilities.
Advanced three-dimensional (3D) mammography systems employing tomography have been developed and are entering the marketplace. In tomography, an x-ray tube is moved through an arc, acquiring many x-ray images over the course of travel. These images may be combined to build up a 3D image aiding cancer diagnosis. There is not, however, a generally accepted practical method or practice for employing the use of conventional anti-scatter grids in 3D tomography-type mammography systems.
Therefore, it would be desirable to design an apparatus and method that overcomes the aforementioned drawbacks.